Method of improving fracture toughness of implantable medical devices through annealing

ABSTRACT

Methods of fabricating a polymeric implantable device with improved fracture toughness through annealing are disclosed herein. A polymeric construct is annealed with no or substantially no crystal growth to increase nucleation density. After the annealing, crystallites are grown around the formed nuclei. An implantable medical device, such as a stent, can be fabricated from the polymer construct after the crystallite growth.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to methods of manufacturing polymeric medicaldevices, in particular, stents.

2. Description of the State of the Art

This invention relates to radially expandable endoprostheses, that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, formed fromwires, tubes, or sheets of material rolled into a cylindrical shape.This scaffolding gets its name because it physically holds open and, ifdesired, expands the wall of the passageway. Typically, stents arecapable of being compressed or crimped onto a catheter so that they canbe delivered to and deployed at a treatment site.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance.Effective concentrations at the treated site require systemic drugadministration which often produces adverse or even toxic side effects.Local delivery is a preferred treatment method because it administerssmaller total medication levels than systemic methods, but concentratesthe drug at a specific site. Local delivery thus produces fewer sideeffects and achieves better results.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesan active or bioactive agent or drug. Polymeric scaffolding may alsoserve as a carrier of an active agent or drug.

The stent must be able to satisfy a number of mechanical requirements.The stent must be capable of withstanding the structural loads, namelyradial compressive forces, imposed on the stent as it supports the wallsof a vessel. Therefore, a stent must possess adequate radial strength.Radial strength, which is the ability of a stent to resist radialcompressive forces, is due to strength and rigidity around acircumferential direction of the stent. Radial strength and rigidity,therefore, may also be described as, hoop or circumferential strengthand rigidity.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. In addition, the stent must possess sufficientflexibility to allow for crimping, expansion, and cyclic loading.

Some treatments with implantable medical devices require the presence ofthe device only for a limited period of time. Once treatment iscomplete, which may include structural tissue support and/or drugdelivery, it may be desirable for the stent to be removed or disappearfrom the treatment location. One way of having a device disappear may beby fabricating the device in whole or in part from materials that erodeor disintegrate through exposure to conditions within the body. Thus,erodible portions of the device can disappear or substantially disappearfrom the implant region after the treatment regimen is completed. Afterthe process of disintegration has been completed, no portion of thedevice, or an erodible portion of the device will remain. In someembodiments, very negligible traces or residue may be left behind.Stents fabricated from biodegradable, bioabsorbable, and/or bioerodablematerials such as bioabsorbable polymers can be designed to completelyerode only after the clinical need for them has ended.

However, there are potential shortcomings in the use of polymers as amaterial for implantable medical devices, such as stents. There is aneed for a manufacturing process for a stent that addresses suchshortcomings so that a polymeric stent can meet the clinical andmechanical requirements of a stent.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method of makingan implantable medical device comprising: annealing a polymer construct,the polymer construct being in a temperature range that allows nucleiformation within the polymer with no or substantially no growth ofcrystallite around the nuclei during a selected annealing time; growingcrystallites around the nuclei after the annealing time to obtain adesired crystallinity in the polymer construct; and fabricating animplantable medical device from the construct after the crystallitegrowth.

Additional embodiments of the present invention include a method ofmaking a stent comprising: annealing a polymeric tube, the polymerictube being in a temperature range that allows nuclei formation withinthe polymer with no or substantially no growth of crystallite around thenuclei during a selected annealing time; growing crystallites around thenuclei after the annealing time to obtain a desired crystallinity; andfabricating a stent from the tube after the crystallite growth.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

FIG. 2 depicts a schematic plot of the crystal nucleation rate and thecrystal growth rate for a polymer.

FIG. 3A depicts a strut of a polymeric stent fabricated withoutannealing.

FIG. 3B is a schematic microstructure of a section of the strut of FIG.3A.

FIG. 4A depicts a strut of a polymeric stent fabricated with annealing.

FIG. 4B is a schematic microstructure of a section of the strut of FIG.4A.

FIG. 5 depicts an axial cross-section of a polymer tube disposed over amandrel with an inner diameter of the tube the same or substantially thesame as an outer diameter of the mandrel.

FIG. 6A depicts an axial cross-section of a polymer tube disposed over amandrel with an inner diameter of the tube greater than an outerdiameter of the mandrel.

FIG. 6B shows the tube of FIG. 6A tube reduced in diameter due toheating.

FIG. 7A depicts an axial cross-section of a polymeric tube positionedwithin a mold.

FIG. 7B depicts the polymeric tube of FIG. 7A in a radially deformedstate.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of the present invention relate to manufacture ofpolymeric implantable medical devices. In particular, the embodimentsinclude a step of annealing a polymer construct during manufacture toincrease the fracture toughness of a device made from the construct. Themethods described herein are generally applicable to any polymericimplantable medical device. In particular, the methods can be applied totubular implantable medical devices such as self-expandable stents,balloon-expandable stents, and stent-grafts.

A stent may include a pattern or network of interconnecting structuralelements or struts. FIG. 1 depicts a view of a stent 100. In someembodiments, a stent may include a body, backbone, or scaffolding havinga pattern or network of interconnecting structural elements 105. Stent100 may be formed from a tube (not shown). The structural pattern of thedevice can be of virtually any design. The embodiments disclosed hereinare not limited to stents or to the stent pattern illustrated in FIG. 1.The embodiments are easily applicable to other patterns and otherdevices. The variations in the structure of patterns are virtuallyunlimited. A stent such as stent 100 may be fabricated from a tube byforming a pattern with a technique such as laser cutting or chemicaletching.

A stent such as stent 100 may be fabricated from a polymeric tube or asheet by rolling and bonding the sheet to form the tube. A tube or sheetcan be formed by extrusion or injection molding. A stent pattern, suchas the one pictured in FIG. 1, can be formed in a tube or sheet with atechnique such as laser cutting or chemical etching. The stent can thenbe crimped on to a balloon or catheter for delivery into a bodily lumen.

An implantable medical device can be made partially or completely from abiodegradable, bioabsorbable, or biostable polymer. A polymer for use infabricating an implantable medical device can be biostable,bioabsorbable, biodegradable or bioerodable. Biostable refers topolymers that are not biodegradable. The terms biodegradable,bioabsorbable, and bioerodable are used interchangeably and refer topolymers that are capable of being completely degraded and/or erodedwhen exposed to bodily fluids such as blood and can be graduallyresorbed, absorbed, and/or eliminated by the body. The processes ofbreaking down and absorption of the polymer can be caused by, forexample, hydrolysis and metabolic processes.

A stent made from a biodegradable polymer is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished. Afterthe process of degradation, erosion, absorption, and/or resorption hasbeen completed, no portion of the biodegradable stent, or abiodegradable portion of the stent will remain. In some embodiments,very negligible traces or residue may be left behind.

The duration of a treatment period depends on the bodily disorder thatis being treated. In treatments of coronary heart disease involving useof stents in diseased vessels, the duration can be in a range from abouta month to a few years. However, the duration is typically tip to aboutsix months, twelve months, eighteen months, or two years. In somesituations, the treatment period can extend beyond two years.

As indicated above, a stent has certain mechanical requirements such ashigh radial strength, high modulus, and high fracture toughness. A stentthat meets such requirements greatly facilitates the delivery,deployment, and treatment of a diseased vessel. With respect to radialstrength, a stent must have sufficient radial strength to withstandstructural loads, namely radial compressive forces, imposed on the stentas it supports the walls of a vessel. In addition, the stent mustpossess sufficient flexibility to allow for crimping, expansion, andcyclic loading. A polymeric stent with inadequate radial strength canresult in mechanical failure or recoil inward after implantation into avessel.

The strength to weight ratio of polymers is smaller than that of metals.To compensate for this, a polymeric stent can require significantlythicker struts than a metallic stent, which results in an undesirablylarge profile. One way of addressing the strength deficiency of polymersis to fabricate a stent from a deformed polymer construct. Deformingpolymers tends to increase the strength along the direction ofdeformation. Thus, a stent fabrication process can include radiallydeforming a polymer tube and cutting a stent from the deformed tube.

With respect to toughness, a polymer stent should also have a highresistance to fracture. Semicrystalline polymers such as poly(L-lactide)(PLLA) that are suitable as stent materials tend to be brittle underbiological conditions or conditions within a human body. Specifically,such polymers can have a glass transition temperature (Tg) above humanbody temperature which is approximately 37° C. These polymer systemsexhibit a brittle fracture mechanism in which there is little or noplastic deformation prior to failure. As a result, a stent fabricatedfrom such polymers can have insufficient toughness for the range of useof a stent. In particular, it is important for a stent to be resistantto fracture throughout the range of use of a stent, i.e., crimping,delivery, deployment, and during a desired treatment period.

A number of strategies may be employed to improve the fracture toughnessof semicrystalline polymers such as PLLA. For example, a rubbery phase(or toughening agent) may be incorporated in the rigid polymer, such aspolycaprolactone or polytrimethylcarbonate through chemical reaction orphysical blending. However, this results in decreased strength andmodulus.

Alternatively, fracture toughness can be improved by reducing the sizeof the polymer crystals and increasing the density of the nuclei fromwhich the crystals grow. Nucleation density can be increased through theaddition of nucleating agents [e.g.,ethylenebis(12-hydroxystearylamide), cyclohexanedicarboxylic dianilide,and tetramethylenedicarboxylic disalicyloylhydrazide]. However, the poorchemical compatibility between these small molecule nucleating agentsand PLLA would engender poor mechanical properties.

Generally, in the crystallization of polymers, there are two separateevents that occur. The first event is the formation of nuclei in thepolymer matrix. The second event is growth of the crystallite aroundthese nuclei. The overall rate of crystallization of the polymer isdependent, therefore, on the equilibrium concentration of nuclei in thepolymer matrix, and on the rate of growth of crystallites around thesenuclei.

Semicrystalline polymers can contain both amorphous and crystallinedomains at temperatures below the melting point. Amorphous regions arethose in which polymer chains are in relatively disorderedconfigurations. Crystalline domains or crystallites are those in whichpolymer chains are in ordered configurations with segments of polymerchains essentially parallel to one another.

The classical view of polymer crystallization is a thermodynamically“frustrated” nucleation and growth process. The transition from thedisordered rubber-like state where flexible chains adopt the random coilconformation to a rigid, ordered, three-dimensional state has beenformally treated as a classical first-order transition. Crystallitesform at the stable nuclei and grow by reorganizing random coil chainsinto chain-folded crystalline lamellae (ca. 10 nm thick). However,individual segments of polymer molecules are often unable to adopt thethermodynamically desirable conformation state necessary forcrystallization before adjacent segments crystallize, locking innon-equilibrium amorphous stricture. Thus, semicrystalline polymers forma mixture of ordered crystalline and disordered amorphous regions. Thecrystalline lamellae form sheaf-like stacks a few lamellae thick (˜50 to100 nm) that splay and branch as they grow outward, forming spherulitesvarying from submicron to millimeters in size. The growth of anindividual spherulite ceases when it impinges with neighboringspherulites. Only in the theoretical limit of infinite time at theequilibrium melting temperature could a semicrystalline polymer form thethermodynamic ideal single-crystal structure.

Hence, for all practical situations, semicrystalline polymers assume akinetically-driven, non-equilibrium morphology in the solid state. Theoverall crystallization kinetics follows the general mathematicalformulation that has been developed for the kinetics of phase changeswith only minor modifications. The importance of nucleation processes inpolymer crystallization has been amply recognized and is based on verygeneral considerations. This concept has been applied to the analysis ofthe kinetics of polymer crystallization.

In general, crystallization tends to occur in a polymer at temperaturesbetween Tg and Tm of the polymer. FIG. 2 shows a schematic of thedependence of nucleation rate (A) and crystal growth rate (B) ontemperature between the glass transition temperature (Tg) and themelting temperature (Tm) under quiescent conditions. At temperaturesabove Tg but far below Tm where polymer chain mobility is limited,nucleation is substantially favored over growth, since the latterprocess requires much more extensive chain mobility. These nuclei remainpresent in the polymer until its temperature is elevated above Tm for aperiod of time. A consequence of the behavior illustrated in FIG. 2 isthat at high temperatures there are relatively few, large crystallitesformed, while at low temperatures, there are relatively more numerous,smaller crystallites formed.

Various embodiments of the present invention of making a polymericdevice can include a step of annealing a polymer construct with no orsubstantially no crystal growth to increase nucleation density. Themethod further includes a step of growing crystallites around the formednuclei after the annealing step. As described in more detail below, thecrystallite growth step can be performed by increasing the temperatureof the construct, deforming the construct, or both. After the crystalgrowth step, an implantable medical device, such as a stent, can befabricated from the polymer construct.

FIG. 3A depicts a strut 100 of a polymeric stent fabricated withoutannealing and FIG. 3B is a schematic microstructure 104 of a section 102of strut 100 showing a small amount of large crystals 106 dispersedwithin an amorphous region 108. FIG. 4A depicts strut 110 of a polymericstent fabricated with annealing. FIG. 4B depicts the schematicmicrostructure 114 of a section 112 of strut 110 showing a large amountof smaller crystals 116 dispersed within an amorphous region 118.

A polymer construct can be a polymer or polymer material formed into ageometrical shape, such as a tube or a sheet. The shape is chosen sothat further processing can be applied to form an implantable medicaldevice. For example, a stent pattern can be cut into a tube to form astent. The polymer construct can be formed using extrusion or injectionmolding. Alternatively, a polymer tube may be formed from a sheet thatis rolled and bonded into a tube.

In such embodiments, the annealing step can include annealing a polymerconstruct at a temperature or temperature range for a selected annealingtime that allows nuclei formation within the polymer with no orsubstantially no growth of crystallite around the nuclei. The annealingseeds nuclei throughout the polymer construct. The temperature range ofannealing can be between Tg and 3, 5, 7, 10, 12, 15 or 18° C. above Tg.Alternatively, the temperature range can be Tg to 0.15×(Tm−Tg).

Exemplary semicrystalline polymers that may be used in embodiments ofthe present invention include PLLA, poly(D-lactide) (PDLA),polyglycolide (PGA), (poly(L-lactide-co-glycolide) (PLGA), andPLLA-b-poly(ethylene oxide) (PLLA-b-PEO). Literature values of ranges ofTg and Tm of PLLA and PGA are given in Table 1.

TABLE 1 Tg and Tm for PLLA and PGA. Melting Glass Transition PolymerPoint (° C.)¹ Temp (° C.)¹ PGA 225-230 35-40 PLLA 173-178 60-65 ¹MedicalPlastics and Biomaterials Magazine, March 1998.

The annealing time can be up to 5 min, 10 min, 30 min, 1 hr, or greaterthan 1 hr. The annealing time can be selected to obtain a desirednucleation density.

As indicated above, after the annealing time, the method furtherincludes a crystal growth step of growing crystallites around the nucleito obtain a desired crystallinity in the polymer construct. A desiredcrystallinity may be at least 10%, 10-20%, 20-30%, 30-50%, or greaterthan 50%. In some embodiments, crystallites can be grown around theformed nuclei by increasing the temperature of the construct to atemperature below Tm that allows crystallite growth. Crystallites can beallowed to grow for a selected crystallite growth time at the increasedtemperature. Additional nuclei can form and crystallites can grow aroundthe additional nuclei at the increased temperature. The temperaturerange for crystallite growth can be any temperature between Tg and Tmthat allows crystallite growth during a selected time period such as upto 1 min, 10 min, 30 min, 1 hr, or greater than 1 hr. For example, thetemperature range can be at least Tg+Z×(Tm−Tg), where Z is 0.1, 0.2-0.4,0.4-0.8, or 0.8-1.

After the crystallite growth step, the construct can then be subjectedto further processing steps in the device fabrication process. Forexample, a stent pattern can be cut into the tube. Alternatively, thepolymer tube can be deformed to increase the strength (as describedbelow) prior to cutting the pattern.

As indicated above, it may be desirable to include a deformation step inthe manufacturing process of an implantable medical device to increasethe strength along the direction of deformation. Such a deformation stepcan also grow crystallites around nuclei formed during the annealingstep. Thus, in some embodiments, the polymer construct is deformed afterthe selected annealing time. In such embodiments, the deformation causescrystalline growth around the formed nuclei. In particular, a tubularpolymer construct can be radially deformed using known methods such asblow molding, that is described below.

In certain embodiments, the temperature of the polymer construct duringdeformation can be higher than the annealing temperature range, butlower than Tm. In such embodiments, growth of crystallites can be due toboth the deformation and the increase in temperature. The temperaturerange of the polymer construct during deformation can be Tg+Y×(Tm−Tg),where Y can be 0.1-0.2, 0.2-0.4, or 0.4-0.8.

In other embodiments, the temperature of the polymer construct duringdeformation can be the same as the temperature during the annealingstep. The deformation process can induce growth of crystallites aroundthe nuclei formed during the annealing step. Growth of crystallitesduring deformation can occur even at temperatures at which there islittle or no crystallite growth at quiescent conditions. As statedabove, the schematic curve (B) for the crystal growth rate in FIG. 2corresponds to quiescent conditions, and, thus, does not apply to thecrystallite growth during deformation. The temperature of the polymerconstruct is desirably above Tg during deformation since as describedbelow, Tg represents a transition from a vitreous state to a soliddeformable or ductile state. Therefore, a temperature above Tgfacilitates deformation of the polymer.

In still further embodiments, a temperature induced crystallite growthstep and a deformation step can be performed sequentially. For example,the temperature can be increased to grow crystallites, followed by adeformation step at a selected temperature. Alternatively, a deformationstep can be performed, followed by equilibrating the deformed constructat an increased temperature that allows crystallites to grow.

Heating and maintaining a temperature of a polymer construct at anannealing temperature or a crystallite growth temperature can beperformed by various methods. For example, the construct in can beheated in a vacuum oven. Alternatively, a warm gas such as nitrogen,oxygen, air, argon, or other gas can be blown on the construct. Thetemperature of the construct can be maintained by known control methods.

A polymer construct may have a tendency to change shape upon heating. Inparticular a polymeric tube may tend to reduce in diameter or shrinkupon heating. In some embodiments, the reduction in diameter of apolymer tube during the annealing step or temperature-induced crystalgrowth steps can be reduced or prevented. Reduction in diameter can bereduced or prevented by disposing a polymeric tube over a mandrel duringthe heating. The shrinkage of the tube is limited to the outsidediameter of the tube. To prevent reduction in diameter, the insidediameter of the tube can be the same or substantially the same as theoutside diameter of the mandrel. FIG. 5 illustrates this with an axialcross-section of a polymer tube 120 disposed over a mandrel 122. Aninner diameter of tube 120 is the same or substantially the same as anouter diameter Dm of mandrel 122.

To reduce shrinkage, the mandrel has an outside diameter less than theinside diameter of the polymer tube. FIG. 6A depicts this with an axialcross-section of a polymer tube 130 disposed over a mandrel 132. Aninner diameter Dt of tube 130 is greater than an outer diameter Dm ofmandrel 132. FIG. 6B shows that as tube 130 is heated during annealingor crystallite growth, tube 130 can reduce in diameter, but that thereduction in diameter is limited to the outer diameter Dm of themandrel.

In further embodiments, shrinkage can be reduced or prevented bymaintaining an increased pressure within the tube. For example, thepolymer tube can be disposed in a mold, e.g., glass, and the internalpressure is increased during heating by blowing a gas in the tube.

As described above, a polymeric tube can be radially deformed using blowmolding. FIGS. 7A-B illustrate an embodiment of deforming a polymerictube using blow molding. FIG. 7A depicts an axial cross-section of apolymeric tube 150 with an outside diameter 155 positioned within a mold160. Mold 160 limits the radial deformation of polymeric tube 150 to adiameter 165, the inside diameter of mold 160. Polymer tube 150 may beclosed at a distal end 170 which may be open in subsequent manufacturingsteps. A fluid is conveyed, as indicated by an arrow 175, into an openproximal end 180 of polymeric tube 150. A tensile force 195 can beapplied at proximal end 180 and a distal end 170.

Polymeric tube 150 may be heated by heating the fluid to a temperatureabove ambient temperature prior to conveying the gas into polymeric tube150. Alternatively, the polymeric tube may be heated by heating theexterior of mold 160 by blowing a warm gas on the mold. The tube mayalso be heated by a heating element in the mold. The increase inpressure inside of polymer tube 150 facilitated by the increase intemperature of the polymeric tube causes radial deformation of polymertube 150, as indicated by an arrow 185. FIG. 7B depicts polymeric tube150 in a deformed state with an outside diameter 190 within mold 160.

Furthermore, the tube may be expanded to a target diameter. In oneembodiment, the target diameter may be the diameter at which a stentpattern is formed by laser machining the tube. The target diameter canalso correspond to the diameter of a stent prior to crimping. The degreeof radial deformation may be quantified by a blow-up ratio or radialdraw ratio:

$\frac{{Outside}\mspace{14mu}{Diameter}\mspace{14mu}{of}\mspace{14mu}{Deformed}\mspace{14mu}{Tube}}{{Original}\mspace{14mu}{Outside}\mspace{14mu}{Diameter}\mspace{14mu}{of}\mspace{14mu}{Tube}}$In some embodiments, the radial draw ratio of a polymeric tube for usein fabricating a stent may be between about 1 and 10, or more narrowlybetween about 2 and 6. Similarly, the degree of axial deformation may bequantified by an axial draw ratio:

$\frac{{Length}\mspace{14mu}{of}\mspace{14mu}{Deformed}\mspace{14mu}{Tube}}{{Original}\mspace{14mu}{Length}\mspace{14mu}{of}\mspace{14mu}{Tube}}$

For the purposes of the present invention, the following terms anddefinitions apply:

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. Tensile stress, for example, is a normal component ofstress applied that leads to expansion (increase in length). Inaddition, compressive stress is a normal component of stress applied tomaterials resulting in their compaction (decrease in length). Stress mayresult in deformation of a material, which refers to a change in length.“Expansion” or “compression” may be defined as the increase or decreasein length of a sample of material when the sample is subjected tostress.

“Strain” refers to the amount of expansion or compression that occurs ina material at a given stress or load. Strain may be expressed as afraction or percentage of the original length, i.e., the change inlength divided by the original length. Strain, therefore, is positivefor expansion and negative for compression.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Modulus” may be defined as the ratio of a component of stress or forceper unit area applied to a material divided by the strain along an axisof applied force that results from the applied force. For example, amaterial has both a tensile and a compressive modulus.

The tensile stress on a material may be increased until it reaches a“tensile strength” which refers to the maximum tensile stress which amaterial will withstand prior to fracture. The ultimate tensile strengthis calculated from the maximum load applied during a test divided by theoriginal cross-sectional area. Similarly, “compressive strength” is thecapacity of a material to withstand axially directed pushing forces.When the limit of compressive strength is reached, a material iscrushed.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness in this caseare in energy per unit volume of material. See, e.g., L. H. Van Vlack,“Elements of Materials Science and Engineering,” pp. 270-271,Addison-Wesley (Reading, Pa., 1989).

The underlying structure or substrate of an implantable medical device,such as a stent can be completely or at least in part made from abiodegradable polymer or combination of biodegradable polymers, abiostable polymer or combination of biostable polymers, or a combinationof biodegradable and biostable polymers. Additionally, a polymer-basedcoating for a surface of a device can be a biodegradable polymer orcombination of biodegradable polymers, a biostable polymer orcombination of biostable polymers, or a combination of biodegradable andbiostable polymers.

It is understood that after the process of degradation, erosion,absorption, and/or resorption has been completed, no part of the stentwill remain or in the case of coating applications on a biostablescaffolding, no polymer will remain on the device. In some embodiments,very negligible traces or residue may be left behind. For stents madefrom a biodegradable polymer, the stent is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.

Representative examples of polymers that may be used to fabricate animplantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate),poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,poly(glycolic acid), poly(glycolide), poly(L-lactic acid),poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(caprolactone), poly(trimethylene carbonate), polyester amide,poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters)(e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin,fibrinogen, cellulose, starch, collagen and hyaluronic acid),polylirethanes, silicones, polyesters, polyolefins, polyisobutylene andethylene-alphaolefin copolymers, acrylic polymers and copolymers otherthan polyacrylates, vinyl halide polymers and copolymers (such aspolyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether),polyvinylidene halides (such as polyvinylidene chloride),polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such aspolystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Another type of polymer based on poly(lacticacid) that can be used includes graft copolymers, and block copolymers,such as AB block-copolymers (“diblock-copolymers”) or ABAblock-copolymers (“triblock-copolymers”), or mixtures thereof.

Additional representative examples of polymers that may be especiallywell suited for use in fabricating or coating an implantable medicaldevice include ethylene vinyl alcohol copolymer (commonly known by thegeneric name EVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

EXAMPLE

The example set forth below are for illustrative purposes only and arein no way meant to limit the invention. The following example is givento aid in understanding the invention, but it is to be understood thatthe invention is not limited to the particular example. The Examplebelow is provided by way of illustration only and not by way oflimitation. The parameters and data are not to be construed to limit thescope of the embodiments of the invention.

PLLA Stent Preparation by Increasing PLLA Nuclei Through AnnealingBefore Tubing Expansion

Step 1 (tubing extrusion): PLLA material is extruded in an extruder at200° C. and the tubing is quickly quenched in cold water or othercooling medium. The size of the extruded tubing is set at about 0.02″for inside diameter (ID) and 0.06″ for outside diameter (OD).Step 2 (tubing Annealing): The extruded tubing is annealed at atemperature between 60 to 75° C. for 30 min to 3 h to create a certainamount of PLLA nuclei.Step 3 (tubing expansion): The annealed tubing is placed in a glass moldand expanded at about 200° F. to obtain biaxial orientation and highercrystallinity. Its final ID and OD are set at about 0.12″ and 0.13″,respectively.Step 4 (stent preparation): A stent is cut from the expanded tubingusing a femto-second laser, crimped down to a small size (0.05″) on aballoon catheter, and sterilized by electron beam at a dose of 25 kGy.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1. A method of making an implantable medical device comprising: annealing a polymer construct, the polymer construct being annealed at a temperature in a temperature range above Tg of the polymer construct that allows nuclei formation to increase nucleation density within the polymer with no crystal growth during a selected annealing time of greater than 1 hr; growing crystallites around the nuclei after the annealing time to obtain a desired crystallinity in the polymer construct, wherein the crystallite growth is induced by deforming the polymer construct; and fabricating an implantable medical device from the construct after the crystallite growth.
 2. The method of claim 1, wherein the implantable medical device is a stent.
 3. The method of claim 1, wherein the polymer construct is tubular and the fabricating comprises cutting a stent pattern in the construct to form a stent.
 4. The method of claim 1, wherein the polymer construct is a tube and the deforming the polymer construct comprises radially expanding the polymer tube to provide circumferential orientation in the polymer tube.
 5. The method of claim 4, wherein the implantable medical device is a stent and the stent is fabricated from the radially expanded tube.
 6. The method of claim 1, wherein the polymer of the polymer construct is PLLA.
 7. The method of claim 1, wherein the polymer of the polymer construct is bioabsorbable.
 8. The method of claim 1, wherein the polymer of the polymer construct is selected from the group consisting of PLLA, PGA, and PLGA.
 9. The method of claim 1, wherein a temperature of the annealing step is between Tg and Tg+15° C.
 10. A method of making an implantable medical device comprising: annealing a polymer construct, the polymer construct being annealed at a temperature in a temperature range above Tg of the polymer construct that allows nuclei formation to increase nucleation density within the polymer with no crystal growth during a selected annealing time of greater than 1 hr; growing crystallites around the nuclei after the annealing time to obtain a desired crystallinity in the polymer construct, wherein the crystallite growth is induced by increasing the temperature of the polymer construct to a range that allows the growth of crystallites during a selected time period and allowing the growth of crystallites to the desired crystallinity in the selected time period; and fabricating an implantable medical device from the construct after the crystallite growth.
 11. The method of claim 10, wherein the selected time period of the growth of crystallites is less than 1 hr.
 12. The method of claim 10, wherein the selected time period of the growth of crystallites is greater than 1 hr.
 13. The method of claim 10, wherein the polymer of the polymer construct is PLLA. 